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Non-invasive imaging plays a central role in cardiovascular disease for determining diagnosis, prognosis, and optimizing patient management. Recent experimental studies have demonstrated that monitoring hyperpolarized 13C-labelled tracers with magnetic resonance imaging and spectroscopy (MRI and MRS) offers a new way to investigate the normal and diseased heart, and that the technology may be useful in patients with heart disease. In this review, we show how hyperpolarized 13C-labelled tracers are generated and have been applied experimentally, and outline the methodological advances currently underway to enable translation of hyperpolarized 13C MRI and MRS into the clinic. Using hyperpolarized 13C-labelled metabolites and metabolic MRI and MRS could help assessment of many human cardiovascular diseases, including coronary artery disease, heart failure and metabolic cardiomyopathies. We discuss the clinical areas in which the technology may, in the future, aid in the diagnosis and management of patients with cardiovascular diseases, including dynamic investigations of in vivo metabolism, coronary angiography and quantitative perfusion imaging. It is possible that, in the future, hyperpolarized magnetic resonance will play a major role in clinical cardiology.
Cardiovascular disease (CVD) is associated with high morbidity, mortality, and financial burden to health care services1-3. In the United States, CVD is the leading cause of death in both men and women, accounting for 1 in every 2.9 deaths in 2006, with coronary disease accounting for 1 in every 6 deaths2. Non-invasive cardiac imaging increasingly plays a fundamental role in diagnosing, assessing prognosis, and monitoring of therapy response in cardiovascular disease1, 4, 5. Two-dimensional echocardiography is the most commonly used imaging modality to measure heart function, due to its low cost and widespread accessibility. Computed tomography (CT), single photon emission computed tomography (SPECT), and positron emission tomography (PET) expose patients to ionizing radiation, but have been used successfully for clinical assessment of coronary arteries, myocardial perfusion, and viability, respectively. Cardiovascular magnetic resonance (CMR) applies no ionizing radiation and is now considered the gold standard in assessing cardiac anatomy, function and mass1. CMR has also shown great potential for evaluating perfusion and viability using gadolinium based contrast agents.
MR spectroscopy (MRS) and MR-based molecular imaging methods have shown promise for evaluation of cardiac metabolism. For example, phosphorus-31 MRS assesses high energy phosphate content and energy reserve in the human heart (reviewed in 6). Other implementations of multi-nuclear MR spectroscopy, including oxygen-17, carbon-13, sodium-23 and proton MRS have described measurement of oxygen consumption7, substrate selection and rates of metabolic flux8, post-infarct sodium accumulation9, and lipid accumulation10, respectively, in ex vivo and in vivo experimental models of disease. MR-based molecular imaging of particles labeled with fluorine-19 nuclei has been used to study tracer and drug pharmacokinetics and metabolism11. Combined PET-MRI methods have been demonstrated in pre-clinical and non-cardiac applications, to assess cardiac parameters in an infarct mouse model12 and for structural, functional and molecular imaging of patients with brain tumors13. However, the use of all MR techniques to assess cardiac metabolism in people, and to image myocardial perfusion and viability without gadolinium contrast, has been limited by an intrinsically low sensitivity.
Hyperpolarization, using the dynamic nuclear polarization (DNP) technique, can yield greater than 10,000-fold signal increases in MR-active nuclei14. When used with MR imaging and/or spectroscopy, hyperpolarized 13C-labelled metabolic tracers allow unprecedented real-time visualization of the biochemical pathways of normal and abnormal metabolism15. Alternately, the spatial distribution of hyperpolarized 13C-labelled agents can be imaged to achieve high contrast for perfusion and angiographic applications.
In November 2010, hyperpolarized [1-13C]pyruvate was administered to patients for the first time, with a view towards using metabolic MR imaging to characterize prostate cancer16. No clinical application in patients with cardiovascular disease has been reported so far; however, given the recent studies in cancer and the experimental application in animal models of cardiovascular disease reported over the past few years, a review of the clinical potential of hyperpolarization techniques in cardiology seems timely. In this manuscript we cover: 1) methods for generation of hyperpolarized 13C MR tracers and their experimental use; 2) possible use of hyperpolarized 13C-labelled tracers for human CVD; and 3) recent and future technological advances needed for translation of cardiac hyperpolarized 13C MR into the clinic.
The acquisition of images using MRI relies on the fact that the protons in water and fat act like small bar magnets when placed inside a strong magnetic field. They either align with, or against, the main magnetic field, creating two distinct populations. The difference in the size of these two populations (known as the polarization) is what determines the strength of the MRI signal (Figure 1). At biological temperatures and field strengths attainable in clinical MR scanners, the difference in these populations is small, leading to a low sensitivity. In clinical MRI, this low sensitivity is compensated for by the high concentration of water and fat in the human body. Unfortunately, this is not the case when using other nuclei that are observable with MRI (e.g. 13C, 31P, 19F); limited natural abundance and low in vivo concentrations prevent routine clinical monitoring of such nuclei.
A potential way to overcome the limitations of low sensitivity is to employ techniques of hyperpolarization that can temporarily align all nuclei in the same direction (Figure 1), increasing the polarization, and hence the signal, of a particular compound. Dynamic nuclear polarization (DNP) is one such technique and is of particular interest for metabolic studies, as it has the potential to dramatically increase the sensitivity of MRI to molecules containing 13C nuclei. Hyperpolarization of 13C labeled metabolites via DNP relies on the fact that electrons have a very high level of polarization (almost 100%, i.e. all electrons aligned in the same direction) when at low temperatures (<1.4 K) and in a high magnetic field (typically 3.35 T). The DNP process, shown in Figure 1, transfers the electronic polarization to the 13C-labelled molecule of interest using microwave irradiation. This is achieved by mixing a source of free electrons (called a radical) with the 13C-labeled sample to be hyperpolarized. The mixed sample is then placed in a high magnetic field and rapidly frozen in liquid helium. The temperature is reduced further by vacuum pumping the sample to very low pressures (~1 mbar) to bring the final temperature to approximately 1 K14. Microwave irradiation, at a frequency determined by the properties of the 13C molecule, the radical and the magnetic field strength, is then applied to transfer the polarization from the electrons to the carbon molecules. Depending on the molecule to be polarized, this process typically takes between 30 to 60 minutes and results in a 13C polarization of up to 50% (i.e. twice as many nuclei pointing in one direction) Such polarization levels are much greater than the normal MRI polarization levels of approximately 0.0005%.
Obviously, the biological applications of a molecule at 1 K are limited, and it is therefore necessary to bring the sample to a physiological temperature before use in an in vivo experiment. To achieve this, the hyperpolarized sample is rapidly dissolved by the injection of a heated and pressurized bolus of aqueous solvent. The resulting solution retains a high level of nuclear polarization (~20-40%) and can be formulated to be at physiological temperature and pH for in vivo injection. The polarization produced then steadily decays back to normal level (thermal equilibrium) at a rate dependent on the inherent properties of the molecule under study (typically 1-2 minutes). Thus, a current limitation of the method is that the enhanced signal is only available for a short period of time.
Theoretically, the DNP technique can be applied to a wide range of molecules labeled with 13C or any other MR-active nucleus. However, for metabolic imaging, several criteria must be met for the successful polarization and in vivo detection of the hyperpolarized molecule and its products. This includes limitations on the molecular properties and the speed of uptake and utilization (for a full review see17). Hyperpolarized 13C MR studies have used the molecules [1-13C]pyruvate, [2-13C]pyruvate and 13C-bicarbonate most frequently, because they polarize efficiently, retain hyperpolarization for a relatively long time (time constants of signal decay ~50 s), and are rapidly metabolized by tissues to provide flux measurements through important enzymes in the heart, as shown in Figure 2.
In this review, we summarize work that has used the DNP dissolution process, coupled with 13C MR, to further our understanding of cardiovascular disease. We also speculate on the technology’s clinical potential, highlighting the areas of cardiovascular medicine where its application appears promising. Rapid advances have been made since the advent of the dissolution DNP process, both in terms of technological understanding and biological application, which means that the technique is used by an increasingly large number of pre-clinical research groups, with particular emphasis on oncology and cardiovascular medicine.
The true potential of hyperpolarized 13C MR lies in its translation to clinical research and clinical diagnosis and management. Towards this goal, the ‘first trial in man’ studies were carried out in November 2010 at the University of California at San Francisco, using hyperpolarized [1-13C]pyruvate in the staging of prostate tumors16. Metabolic signatures appropriate to benign and cancerous tissue were seen, with elevated lactate levels within the tumor. Achieving the goal of routine clinical application will require considerable technological advances, in terms of improved methods and hardware for the acquisition of 13C images, and access to affordable hyperpolarization tools and 13C-labelled compounds. Further, the widespread use of hyperpolarized 13C MR methods in the clinic will require the technique to provide a clear advantage over other non-invasive diagnostic technologies in improving clinical outcome.
Hyperpolarized 13C MR offers many theoretical advantages over existing metabolic imaging techniques (e.g. PET). Firstly, the method does not use ionizing radiation, making the procedure an ideal candidate to satisfy the current demand for cardiovascular imaging strategies that minimize patient radiation exposure18, 19. Secondly, the inherent ability of MRI to encode both spatial and spectral information means that it is possible to distinguish between the injected tracer and its downstream metabolic products. This enables metabolic pathways to be tracked over multiple enzyme-regulated steps, potentially yielding sensitive and specific indicators of disease. Thirdly, hyperpolarized 13C-labelled tracers are normal physiologically occurring compounds, minimizing the risk of adverse effects from pharmacological interactions and potentially providing an alternative to contrast agents that are contraindicated in certain patient groups (e.g. gadolinium-based agents in patients with advanced kidney disease). Finally, the potentially high cost of hyperpolarized 13C metabolic tests can be minimized by incorporating the hardware into existing MR facilities and the rapid 13C MRI and MRS scans into standard CMR protocols. Such incorporation should be relatively trivial as the additional hardware demands (multinuclear transmit/receive channels and RF coils) are simple additions to existing MRI equipment.
All imaging techniques have their specific disadvantages and hyperpolarized 13C MR is no exception. The greatest challenge presented by the technique is the rapid decay of MR signal following DNP and dissolution, which means that a hyperpolarized tracer must be injected in vivo as soon as possible after it is produced. Signal decay restricts the compounds that can be hyperpolarized: their relaxation must be slow enough to ensure that they can be administered in vivo and, once administered, they must be metabolized quickly. Additionally, to minimize tracer delivery time, the polarizer equipment must be sited immediately beside the MR system, which introduces the need for sterile handling of compounds within the MRI suite, and the associated regulatory issues. This issue has been addressed to some extent by the design of a sterile polarizer20 and the granting of an ‘Investigational New Drug’ approval for the use of hyperpolarized [1-13C]pyruvate by the FDA16. However, the regulatory issues surrounding the widespread use of new hyperpolarized compounds remain a hurdle to be addressed.
A second limitation of the technique is that, despite the large gains in sensitivity over thermal equilibrium MRI, the technique does not reach the sensitivity of PET, which can detect extremely low concentrations of radioactive tracer in the nM range21, albeit with a low spatial resolution. Therefore, acquisition of sufficient MR signal requires the injection of metabolic tracers at doses in the μM - mM range. This disadvantage is partially ameliorated by the fact that the infused compounds are naturally occurring. However, metabolic tracer levels that approach or exceed physiological levels of the unlabelled compound will be required, and need to be shown to be safe in humans. Thus the metabolic effects of the injection itself will need to be well understood when interpreting results.
To meet its task of continually circulating blood throughout the body, the heart consumes more energy in the form of ATP than any other organ6. The healthy heart derives 60-90% of its energy from the oxidation of fatty acids, with the remainder primarily from pyruvate oxidation, derived from glucose (via glycolysis) and lactate. However, when plasma substrate composition is altered, the relative contributions of lipids, carbohydrates, and ketone bodies to cardiac ATP production vary substantially22, 23. Before oxidation, all substrates are converted to acetyl-CoA, a two-carbon molecule that enters the Krebs cycle. Carbohydrate-derived pyruvate is transported into the mitochondria and irreversibly converted to acetyl-CoA by the pyruvate dehydrogenase enzyme complex (PDH). Fatty acids are converted to acetyl-CoA via mitochondrial β-oxidation. Randle described the quantitative competition for respiration between glucose and fatty acids, and the variation in their roles as major fuels for the heart, as the glucose–fatty acid cycle24.
In vivo control of the PDH enzyme complex is a fundamental determinant of the relative contributions of glucose and fatty acid oxidation to ATP production in the heart22, 23, 25. Several PDH-mediated mechanisms exist to promote fatty acid oxidation over glucose oxidation, including phosphorylation-inhibition of PDH26 and end-product inhibition of PDH by acetyl-CoA and NADH27. However, metabolic flexibility remains the key characteristic of the heart to meet high ATP demand 28, 29. After eating and in response to energetic stressors, such as aerobic exercise or ischemia, glucose and other carbohydrates become more predominant cardiac fuels 30, 31.
Initially, hyperpolarized 13C MRS measurements of in vivo substrate selection were validated via comparison with analogous data collected in vitro and ex vivo. It was first demonstrated in the isolated perfused rat heart that infusion of hyperpolarized [1-13C]pyruvate, and MRS detection of total carbonic acid (13CO2 plus 13C-bicarbonate), measured flux through the PDH enzyme complex32. The operation of the glucose-fatty acid cycle in vivo was subsequently demonstrated in a study using hyperpolarized [1-13C]pyruvate and MRS to measure PDH flux in fed and fasted rats (Figure 3A). As Randle observed in vitro in 196324, elevated plasma free fatty acids, typical of the fasted state, reduced PDH-mediated glucose oxidation in vivo, compared with the fed state33. The direct correlation between PDH flux measured in vivo using hyperpolarized [1-13C]pyruvate and PDH activity measured using an in vitro assay has been demonstrated (Figure 3B) in rats in which PDH activity was increased from low to high rates (0.77 to 6.75 μmol.min.gww) using metabolic interventions including high-fat feeding and pharmacological PDH activation with dichloroacetate34. Thus, the study of PDH activity and its control has been a major application of hyperpolarized [1-13C]pyruvate.
Real-time Krebs cycle metabolism has also been followed in the heart using 13C MRS, by shifting the 13C-label to the second carbon of pyruvate35. When hyperpolarized [2-13C]pyruvate is used as a metabolic tracer, the 13C-label is retained within acetyl-CoA rather than being released as 13CO2, enabling downstream metabolic steps to be visualized (Figure 4). The enzymatic conversions of hyperpolarized [2-13C]pyruvate to [2-13C]lactate, [1-13C]acetylcarnitine, [1-13C]citrate and [5-13C]glutamate were observed with sub-second temporal resolution in the perfused and in vivo rat heart35, 36. The appearance of 13C in the glutamate pool was delayed by 3 s compared with citrate in vivo, demonstrating the feasibility of measuring instantaneous flux through the first span of the Krebs cycle and the oxoglutarate-malate carrier (OMC).
Rapid mitochondrial cycling between acetyl-CoA and an acetylcarnitine substrate buffer was revealed by following real-time hyperpolarized [2-13C]pyruvate metabolism36. Understanding the role of acetylcarnitine in ‘fine-tuning’ mitochondrial acetyl-CoA supply is important in the healthy heart, as carnitine deficiency has been reported in pathophysiological conditions in which cardiac energetics are also depleted, including old age37, pressure-overload hypertrophy38, and heart failure39, 40. Further advances to our understanding of cardiac metabolism will undoubtedly be forthcoming, as for the first time hyperpolarized 13C MR enables in vivo metabolic processes to be monitored, implying that results will be obtained with 1) preserved myocyte structure and environment, 2) physiologically high cardiac workload, and 3) maintenance of the plasma substrate and neurohumoral composition (with the exception of minor disturbances caused by the metabolic tracer itself).
In future, when hyperpolarized 13C MR methods are approved for cardiovascular use in humans, studies of metabolism in the healthy heart will need to be performed. These studies will be vital to optimizing protocols for cardiac hyperpolarized 13C MRI and MRS, in terms of pre-scan preparation to establish a reproducible metabolic state, identification of tracer doses required, and refinement of MR data acquisition and data processing methods. In addition, defining metabolic substrate selection non-invasively in the healthy human heart will be fundamental to the development of drugs targeting cardiac metabolism. In particular, showing that the glucose-fatty acid cycle operates in humans is of great importance; excess free fatty acid oxidation at the expense of glucose oxidation may adversely affect the heart in a variety of conditions, caused by raised plasma free fatty acids levels and causing the insulin resistance found in diabetes, the metabolic syndrome, obesity, and chronic heart failure6, 41, 42.
There is preliminary evidence that hyperpolarized 13C MRI could have a role in diagnosing ischemic heart disease by directly detecting the metabolic consequences of myocardial ischemia, assessing the coronary arteries, mapping myocardial perfusion and demonstrating myocardial viability.
The metabolic consequences of ischemia have been extensively characterized in experimental animal models. In ischemia, primary metabolic adaptations include 1) increased reliance on glycogen breakdown and glucose uptake as a source of energetic fuel43; and 2) reliance on anaerobic glycolysis for energy production44, with decreased Krebs cycle metabolism and oxidative phosphorylation due to low oxygen availability45. The end products of glycolysis include protons (with an associated acidosis), lactate, and NADH that increases the reduction-oxidation (redox) potential of affected cardiomyocytes43, 44, 46-48.
Experiments performed in the globally ischemic, isolated rat heart have demonstrated that hyperpolarized 13C-pyruvate can show the glycolytic switch characteristic of myocardial ischemia. In hearts perfused with Krebs-Henseleit buffer and either pyruvate alone49 or glucose plus pyruvate35, 50, total global ischemia for 10 min resulted in substantially increased 13C-lactate levels compared with controls. In addition, a study using hyperpolarized [2-13C]pyruvate demonstrated reduced oxidative metabolism immediately upon reperfusion, as indicated by reduced flux of the 13C-label into the [1-13C]citrate and [5-13C]glutamate pools35.
Hyperpolarized tracers also enable non-invasive measurement of intracellular and extracellular pH (pHi and pHe). After infusing hyperpolarized 13C-bicarbonate intravenously in mice with subcutaneously implanted tumors51, MR images showed the distribution of hyperpolarized 13C-bicarbonate and 13CO2. A pHe map was then generated throughout the tumor using the 13C-bicarbonate/13CO2 ratio and the Henderson-Hasselbalch equation (Figure 5). In tumors, this approach succeeded because the high carbonic anhydrase activity on the surface of tumor cells52 and within erythrocytes53 ensured virtually instantaneous equilibrium between the infused 13C-bicarbonate and 13CO2. In the heart, use of hyperpolarized 13C-bicarbonate would also be expected to accurately measure pHe in vivo, as the red cell carbonic anhydrase activity is maintained. This may become a sensitive imaging marker of ischemia.
Hyperpolarized 13CO2 and 13C-bicarbonate generated as a byproduct of [1-13C]pyruvate metabolism can measure in vivo cardiac pHi50. Isolated perfused heart studies, in vivo studies and mathematical modeling established that 13CO2 produced by mitochondrial PDH remains within cardiomyocytes for a sufficiently long time to equilibrate with 13C-bicarbonate. Therefore, detection of the cardiac 13C-bicarbonate/13CO2 ratio after hyperpolarized [1-13C]pyruvate infusion offers a novel non-invasive technique to measure cardiac pHi. Magnetization transfer 31P MRS54 and Amide Proton Transfer Imaging55 have previously been used to measure pHi in vivo; however low sensitivity and limited field homogeneity due to cardiac motion, respectively, imply that hyperpolarized [1-13C]pyruvate may be the first method for mapping cardiac pHi that could realistically be applied in patients.
In CAD patients, myocardial ischemia is a regional phenomenon with variable transmural severity, being more severe in the endocardium compared with the epicardium and ranging in severity from transient and reversible (angina, stunning, hibernation) to prolonged and irreversible (myocardial necrosis and scar). In between these two extremes, many clinical situations can be recognized in which a mixture of viable and necrotic/scarred tissue exists48. Characterizing the metabolic patterns unique to ischemic myocardium in CAD patients has been complicated by tissue heterogeneity and the difficulty of making non-steady-state measurements of lactate production and/or pH from the precise site of the ischemic zone48, 56. Therefore, use of hyperpolarized 13C MRI to image localized changes in lactate accumulation and pH non-invasively, and with high spatial resolution, may be valuable in clinical research to reveal the metabolic changes occurring with acute ischemia/reperfusion injury in humans.
Using hyperpolarized 13C MRI to show ischemia may have a direct clinical application in determining the immediate tissue outcome of procedures, such as percutaneous coronary intervention (PCI) or coronary artery bypass graft (CABG). MR images of hyperpolarized [1-13C]pyruvate metabolism that show normal pHi (~7.0–7.2 46) and minimal 13C-lactate production may indicate that blood flow is adequately restored57, 58. Further, hyperpolarized 13C MRI examination immediately after primary PCI may indicate a patient’s potential for recovery and degree of irreversible damage. Hyperpolarized 13C MRI may also contribute to clarifying the role of ischemia in cardiac syndrome X59.
Despite the associated doses of ionizing radiation, cardiac computed tomography (CT) is the current gold standard for non-invasive assessment of the coronary arteries, based on calcium scoring and CT coronary angiography60. Although progress has been made61, reliable coronary lumen imaging with CMR has been hindered by the high temporal and spatial resolution required to accurately image stenosed arteries1. However, when hyperpolarized 13C-labelled substrates are imaged with MRI, the lack of background signal means that acquired images only show structures containing the infused contrast agent62. The resultant gains in signal-to-noise ratio and contrast-to-noise ratio may enable rapid, high resolution imaging of the coronary arteries using MRI, which would allow direct and repeatable assessment of coronary arteries without the use of ionizing radiation.
Towards this goal, Golman et al. provided the first evidence that hyperpolarized 13C MRI, using either dissolution DNP methods or other hyperpolarization techniques, may be useful to visualize the coronary arteries62. Hyperpolarized 13C-urea was infused via a catheter after selective intubation of the left coronary artery in healthy pigs, and the coronary arteries were imaged using a balanced steady-state free procession (b-SSFP) MRI acquisition sequence. As an initial proof-of-concept study, the results were promising. However, the images were acquired using a projection imaging technique (slice thickness of 15 cm) and the in-plane spatial resolution (~2 mm × 2 mm) was an order of magnitude away from that achievable with CT. The temporal resolution of the images was 422 ms (one image acquired every heart beat), still some way from that achievable with CT (~ 80 ms).
The limited spatial resolution currently achievable with hyperpolarized 13C MRI is the result of both the short acquisition time (caused by the decay of the enhanced signal) and the demands that 13C imaging places on the gradient system of the MRI system due to the low gyromagnetic ratio of 13C (4 times lower than 1H). Future hardware improvements may help to overcome the gradient limitations and the development of new intravascular hyperpolarized agents with a longer relaxation time, possibly based on other nuclei (e.g. 15N), may improve the achievable spatial resolution. However, the need to limit background signal may require direct infusion of the tracer into the coronary arteries. This would still necessitate invasive catheterization, limiting the advantage of hyperpolarized 13C MRI over conventional x-ray coronary angiography to the elimination of the radiation burden.
In the wake of the COURAGE trial63, there has been increased importance placed on demonstrating inducible myocardial ischaemia before initiating an elective coronary intervention. Magnetic resonance imaging of myocardial perfusion with a gadolinium-based contrast agent during pharmacological stress has emerged as a highly sensitive and specific technique for non-invasive demonstration of reversible, stress-inducible ischemia64. MRI has a number of advantages over other established techniques, such as single photon emission computerized tomography (SPECT) and stress echocardiography, including high spatial and temporal resolution, no exposure to ionizing radiation, no attenuation or scatter artifacts and no image orientation constraints. However, quantitative analysis of MR perfusion images remains time consuming, and most centers continue to limit analysis to visual assessment64, 65. This is usually sufficient to identify perfusion defects, though quantitative image analysis may be needed to identify stress-inducible ischemia in patients with 3-vessel disease and balanced myocardial hypoperfusion64.
Unlike gadolinium-based contrast agents, hyperpolarized agents directly generate MR signal, so that 13C MR signal intensity is proportional to tissue 13C concentration and therefore tissue perfusion62. Imaging the distribution of a non-metabolized hyperpolarized agent is a theoretically simple method for quantitative perfusion mapping to demonstrate stress-inducible ischemia. Golman et al. have demonstrated the feasibility of imaging myocardial 13C distribution to map perfusion. Hyperpolarized 13C-urea was infused intravenously into a pig, and images were acquired with a b-SSFP pulse sequence and 3 mm × 3 mm in-plane spatial resolution with 10 mm slice thickness62. Upon LAD occlusion, no signal was measured in the corresponding coronary territory. Further, the authors calculated quantitative myocardial perfusion maps on a pixel-by-pixel basis using the Kety-Schmidt method66, after correcting for hyperpolarized tracer signal decay.
To extend hyperpolarized 13C MRI methods of quantitative perfusion imaging to patients, hyperpolarized agents that are not metabolized in vivo should be developed that have long relaxation times, and their ability to indicate myocardial perfusion should be evaluated under rest and adenosine stress conditions. Furthermore, images illustrating both the perfusion and metabolism of hyperpolarized [1-13C]pyruvate at rest and during adenosine stress should be compared. In stress-inducible ischemia, it may be possible to detect regional increases in the conversion of [1-13C]pyruvate into [1-13C]lactate, which could indicate perfusion defects in individual coronary territories.
Incorporation of hyperpolarized agents into conventional MR myocardial perfusion imaging methods could generate a technique that is, conceptually, much like quantitative SPECT but without the use of ionizing radiation. Furthermore, the spatial resolution achieved by Golman et al (3 mm in-plane, 10 mm slice thickness) is already superior to the spatial resolution attainable with SPECT, and the future development of novel pulse sequences for 13C perfusion mapping will undoubtedly improve this resolution further.
While hyperpolarized 13C-labeled agents may simplify quantification relative to conventional MR based perfusion methods, their use in the clinic would increase imaging complexity and cost, and it is currently unclear if the spatial and temporal resolution of hyperpolarized 13C MRI will be able to compete with that of gadolinium contrast agent-based perfusion methods. Areas of potential advantage for the use of hyperpolarized 13C MRI over gadolinium based approaches may be to demonstrate stress-inducible ischemia in patients with suspected 3-vessel disease and balanced hypoperfusion, who require quantitative perfusion imaging (which should be more straight forward for hyperpolarized approaches), and patients with advanced kidney failure, who would not be suitable for gadolinium-based contrast agents.
In patients with ischemic heart disease, surgical revascularisation of akinetic, yet viable, tissue significantly improves outcome67-71. The use of non-invasive imaging to identify regions of stunned and hibernating versus necrotic and scarred myocardium (i.e. the determination of tissue viability) is essential for the optimal treatment of patients with ischemic heart disease and reduced LV function considered for revascularization67, 68, 71-73. Currently, the most promising techniques to assess myocardial viability include CMR with late gadolinium enhancement (LGE) and PET perfusion/metabolism mismatch studies performed using the tracers 18F-FDG and 13N-ammonia74. In CMR with LGE, viability can be quantified by the transmurality of hyperenhancement in dyskinetic or akinetic segments1, 70, and the transmural extent of scar correlates closely with the likelihood of segmental recovery70, 75.
A study by Golman et al used CMR with LGE as a gold standard to verify that hyperpolarized 13C MR could demonstrate myocardial viability62, 76. Hyperpolarized [1-13C]pyruvate was infused intravenously into a pig, following either a 15 min or a 45 min occlusion of the LCx and 2 hours of reperfusion. 13C metabolic images were acquired with an in-plane spatial resolution of 5 mm × 7.5 mm, slice thickness of 20 mm and a total acquisition time of 13.4 s. Metabolic images were compared with CMR-LGE images acquired 15 min after gadolinium infusion, with 3 mm × 3 mm in-plane spatial resolution and a slice thickness of 10 mm. Following the 15 min occlusion, CMR revealed hypokinetic wall motion with no accompanying hyperenhancement on LGE, indicative of stunned myocardium. The accompanying metabolite maps revealed suppressed PDH-mediated 13C-bicarbonate production that was co-localized with the hypokinetic areas (Figure 6). Following a 45 min occlusion, hyperenhancement indicating irreversible damage was detected, and both 13C-bicarbonate and [1-13C]alanine production were reduced in the corresponding area (Figure 6). This suggests that not only can hyperpolarized 13C distinguish between viable and necrotic myocardium, as can LGE, but, unlike LGE, it can also distinguish between normal and akinetic viable myocardium, due to the unique metabolic signatures of each condition.
One major advantage of hyperpolarized methods over CMR-LGE and PET may be the time-efficiency of the method. In the Golman study, hyperpolarized 13C MRI only required one bolus of contrast agent and a 13.4 s scan time to distinguish between viable and non-viable tissue. Clinically this would provide a logistic benefit over CMR-LGE, which requires scan times of 5-20 min, and PET, which requires a dual bolus of radioactive tracer and scan times approaching 1 h77.
Future development of the technique will be necessary to determine if hyperpolarized 13C MRI can assess myocardial viability at relevant doses of infused [1-13C]pyruvate, and if the technique will bring the spatial resolution required to assess tissue in heterogeneous ischemic zones and to determine the transmural extent of scar. Although the spatial resolution of hyperpolarized 13C MRI compares favorably to that of PET, it remains to be seen whether improvements to 13C data acquisition methods can achieve spatial resolution that approaches the resolution attainable with CMR-LGE.
The occurrence of metabolic dysfunction in heart failure is undisputed, yet the causal role(s) of altered substrate utilization and myocardial energetics remains controversial6, 23. Patient studies performed using 31P MRS and PET imaging have confirmed that cardiac energetics are abnormal throughout the progression of heart failure. Myocardial phosphocreatine/ATP ratios, as determined using 31P MRS, are reduced in heart failure, correlate with New York Heart Association classes78 and with indices of systolic78 and diastolic79 function, and predict prognosis80. PET studies have demonstrated that, throughout the progression of heart failure, the balance between glucose and fatty acid oxidation shifts81-83, and that many treatments including β-blockade and cardiac resynchronisation therapy normalise myocardial oxidative metabolism and efficiency82, 84, 85. However, controversy remains over the exact nature of these metabolic alterations86, with the relative utilization of fatty acids either increased, maintained or decreased depending on the aetiology and stage of disease6, 23, 87. Changes in glucose utilization in the failing heart are also poorly understood6. This controversy is due, in part, to the fact that myocardial metabolic investigations are either carried out using destructive in vitro methods or using in vivo radio-labelled tracer techniques, including PET. PET studies in patients have yielded inconsistent results describing the nature of shifts in substrate selection that occur with heart failure, depending on the disease aetiology, patient preparation prior to the study, and if the PET tracers were oxidised (11C-glucose and 11C-palmitate81) or measured only substrate uptake (18F-FDG and 18F-fluoro-6-thiaheptadecanoic acid82, 83). To understand the timing and consequences of switches in substrate metabolism during the development of heart failure, and to potentially use those switches for disease diagnosis, a non-invasive method will be required that is capable of serially monitoring in vivo cardiac metabolism86.
Studies are currently using hyperpolarized 13C MR to examine substrate selection throughout the progression of heart failure, in a rat model of LV remodeling following myocardial infarction88, and a porcine tachycardia-induced model of dilated cardiomyopathy89. These studies have serially examined hyperpolarized [1-13C] and [2-13C]pyruvate metabolism with MRS and/or MRI and have demonstrated reduced incorporation of the 13C label into the [5-13C]glutamate pool throughout the progression of the disease.
In future, if serial 13C measurements are correlated with cine-MRI measurements of cardiac structure and function, 31P MRS measurements of cardiac energetics, and biochemistry and histology, it may be possible for these studies to gain unique insight into the temporal relationships that exist between alterations in substrate metabolism, energetics and function, and to determine if altered metabolism is a cause or consequence of heart failure86. Distinct profiles of 13C-labelled tracer metabolism may emerge that specifically correlate with cardiomyopathy stage, genotype and aetiology. If this is the case, metabolic profiling with hyperpolarized 13C MR may become a useful technique for cardiologists to predict heart failure patient prognosis and ability to recover function, and to optimise disease management based on stage and aetiology6, 86, 90. For example, determination of 13C metabolic profiles may help to diagnose hypertrophic cardiomyopathy (HCM) in cases of borderline or overlapping phenotypes.
Applications of hyperpolarized 13C MR to the failing heart may also aid the design and evaluation of novel treatments. New drugs for normalization of cardiac metabolism and energetics may emerge, based on specific metabolic defects identified by hyperpolarized 13C MR. Serial 13C metabolic profiling may also be useful to monitor the efficacy of new heart failure treatments, particularly metabolic modulators such as trimetazidine91, 92 and carnitine palmitoyltransferase-1 inhibitors93, 94, but also including non-metabolic drugs and interventional therapies that may improve cardiac efficiency82, 84, 85. Finally, metabolic profiling with hyperpolarized 13C MR may be used to assess transplanted hearts to detect early signs of rejection.
Diabetic cardiomyopathy is defined as the alteration of cardiac structure and function induced independently by diabetes mellitus in the absence of ischemic heart disease, hypertension or other cardiac pathologies95-97. A consensus has yet to be reached regarding the precise diagnostic criteria of diabetic cardiomyopathy, though its diagnosis currently relies on two important components96: 1) the detection of myocardial abnormalities, usually including evidence of hypertrophy and LV diastolic dysfunction determined via non-invasive echocardiographic or CMR methods, and 2) the exclusion of other contributory causes of cardiomyopathy. The precise pathophysiological mechanisms leading to the development of diabetic cardiomyopathy also remain unclear98. However, the altered myocardial substrate metabolism characteristic of diabetes mellitus has been linked with cardiomyopathy progression97, 99, 100, suggesting that hyperpolarized 13C MR may play a role in clarifying the mechanisms of disease and in developing more specific diagnostic criteria.
In animal models of diabetes, the heart switches almost exclusively to fatty acids and ketone bodies for its ATP requirements, at the expense of glucose and pyruvate oxidation23, 99. This switch in substrate utilization is dominated by decreased PDH activity101-103, due to increased plasma free fatty acids. Hyperpolarized [1-13C]pyruvate with MRS has been used to measure in vivo cardiac substrate selection in rat models of acute streptozotocin-induced type 1 diabetes and insulin resistance induced by 28 days of feeding a high fat diet. In type 1 diabetic rats, PDH flux was 65% lower compared with control rats and correlated with disease severity33, whereas insulin resistant rats had a 56% reduction in PDH flux34. In future, serial studies with hyperpolarized [1-13C]pyruvate, [2-13C]pyruvate and potentially other metabolic tracers that target oxidative metabolism downstream of PDH should be performed to examine longitudinal changes of in vivo metabolism in chronic models of insulin resistance, metabolic syndrome, and diabetes. Collectively these studies could more fully illustrate the metabolic alterations involved in the onset and progression of diabetic cardiomyopathy.
Hyperpolarized 13C MR measurements of cardiac glucose oxidative capacity (at the level of PDH flux) could optimize the use of novel therapies for diabetes and cardiomyopathy that improve glycaemic control and normalize substrate selection, such as glucagon-like peptide analogues104, 105 and dipeptidyl peptidase 4 inhibitors106. Additionally, clinical studies using cardiac hyperpolarized 13C MR methods to serially examine diabetic patients could 1) reveal the role of altered substrate selection in compromising cardiac structure and function; 2) recognize an early-stage metabolic profile indicating that a diabetic patient may be at risk for developing cardiomyopathy; and 3) identify the profile of 13C substrate utilization confirming that a patient’s heart failure is related to diabetes.
Evidence is emerging that many patients with ‘idiopathic’ cardiomyopathy have inherited defects in either mitochondrial energy metabolism or sarcomeric proteins 107. The so-called ‘metabolic cardiomyopathies’ develop in the context of a wide spectrum of systemic pathological conditions. In many cases, patients present with symptoms of systemic disease, though in some conditions, including Anderson-Fabry disease (a disorder of lysosomal metabolism caused by α-galactosidase A deficiency), several types of glycogen storage disorders, inborn errors of fatty acid metabolism and myopathic carnitine deficiency, patients may initially present with cardiomyopathy1, 107-109. Due to the rare occurrence of many metabolic cardiomyopathies and the poor correlation between in vitro enzyme assays and disease severity107, initial diagnosis can be difficult and can often only be made conclusively via genetic evaluation108.
When hyperpolarized 13C MRI and MRS methods are approved for human use, determining the 13C metabolic profiles of patients with metabolic cardiomyopathies could help to clarify the specific biochemical defects that underlie the development of cardiac disease. Further, in patients with systemic metabolic disorders that initially and predominantly manifest as a cardiomyopathy, 13C metabolic profiling could confirm the metabolic basis of disease and suggest the optimal course of treatment. As an example, in vivo metabolism in the hyperthyroid rat heart has been investigated with [1-13C]pyruvate and MRS110. The term “hyperthyroidism” encompasses a heterogeneous group of disorders, all characterized by elevated levels of thyroid hormones and an increased basal metabolic rate. Hyperthyroidism causes many systemic effects which markedly affect the heart, increasing heart rate, contractility and cardiac output, causing hypertrophy, and altering myocardial substrate selection111. Hyperpolarized 13C MRS revealed a 55% reduction in PDH flux in hyperthyroid animals, compared with controls, a result that helped to clarify one cause of thyroid hormone-induced cardiomyopathy110.
It is likely that carnitine deficiency may be detectable by following the conversion of hyperpolarized [2-13C]pyruvate into [1-13C]acetyl-carnitine, and simple carnitine supplementation reverses the symptoms of intractable congestive heart failure within 5 months109. Additionally, glycogen storage disorders and Anderson-Fabry Disease can be misdiagnosed as HCM1, 108, 112, and the different clinical courses associated with HCM and metabolic cardiomyopathies underscore the importance of accurate diagnosis. For example, glycogen storage disease has been linked to progressive conduction-system disease that may necessitate the implantation of a pacemaker and aggressive control of arrhythmias108, whereas patients with Anderson-Fabry disease respond to enzyme-replacement therapy113. Metabolic phenotyping of these conditions with hyperpolarized 13C MRS could make an important clinical contribution to the diagnosis and management of such patients.
Eventual translation of hyperpolarized 13C methods into the clinic will require considerable technological advances, in terms of improved methods and hardware for the acquisition of 13C images. In order to identify focal regions of ischemia using lactate and/or pH measurements, for example, three-dimensional images of [1-13C]lactate, H13CO3− and 13CO2 with high spatial resolution will be required. Recently, a chemical-shift specific, cardiac- and respiratory-gated, multi-slice spiral MR imaging method was demonstrated with the capacity to map the distribution of [1-13C]pyruvate, [1-13C]lactate, and H 13CO3− with 8.8 mm in-plane spatial resolution and 10 mm slice thickness in normal pig hearts in vivo (voxel size of 0.774 ml)114, which suggests the feasibility of acquiring such data from patients (Figure 7). Further, successful dynamic lactate imaging with a voxel size of 0.125 ml has been demonstrated in tumors115, and despite the added challenges of acquiring MR images from the beating heart it is likely that the spatial resolution attainable for dynamic 13C-lactate imaging will improve towards this level. Cardiac metabolic imaging will also benefit from the design and implementation of outer volume suppression RF pulses that ensure metabolites generated outside the heart do not wash into the region of interest and confound studies of cardiac enzymatic fluxes116.
Development of metabolic tracers beyond [1-13C]pyruvate, [2-13C]pyruvate, and 13C-bicarbonate will enable additional metabolic enzymes to be studied in the heart. Distinct spans of the Krebs cycle may be assessed in the heart using either hyperpolarized [1-13C]glutamate or [1,4-13C2]fumarate, whose respective conversions to [1-13C]α-ketoglutarate and [1,4-13C2]malate have been demonstrated both in vitro and in vivo17, 117. Additionally, hyperpolarized [1,4-13C2]fumarate may emerge as a specific metabolic marker of cell death by necrosis117. Hyperpolarized [1-13C]acetate may form the basis of an assay for intracellular CoA levels118 or, like the PET tracer 11C-acetate, may enable measurement of total metabolic capacity84, 85.
A recent proof-of-concept study has demonstrated that a ‘secondary hyperpolarization’ approach, in which DNP-hyperpolarization is transferred catalytically between molecules119, may enable polarization of metabolites that otherwise would not polarize efficiently. This, along with other techniques using long lived ‘singlet states’120, may extend the hyperpolarized lifetime and therefore the potential imaging window. Further, an increase in attainable polarization, from the typical polarization levels of ~30% described previously, to ~60%, has recently been achieved by increasing the magnetic field strength of the polarizing magnet121 and will aid detection of low concentration 13C-labelled metabolites and enable development of new metabolic tracers.
Quantification of instantaneous metabolic fluxes will be essential for future development of both basic science and translational applications of hyperpolarized 13C MR. Basic science studies should strive towards both measuring metabolic fluxes in units that can be compared with conventional biochemical assays (i.e. μmol/min/g), and describing 13C flux through dynamic metabolic pools with rapid metabolic turnover, where 13C may not accumulate, such as [1-13C]citrate and [1-13C]acetyl-CoA. In the clinic, a system for monitoring the effects of the hyperpolarized agent’s input function on metabolic images must be developed, as tracer pharmacokinetics could differ dramatically between control subjects and patients89. A pulse sequence that dynamically images infused hyperpolarized [1-13C]pyruvate with a low flip angle has recently been demonstrated in pigs, and is the first step towards input function quantification122. Finally, reproducibility studies to identify quantitative biomarkers that are both sensitive and specific to cardiovascular disease must be performed. Towards this goal, ratiometric markers that self-normalize to basal variations in tracer pharmacokinetics and metabolism may prove useful.
In conclusion, when used with MRI and MRS, hyperpolarized 13C-labelled tracers offer the first method to measure cardiac substrate metabolism in real-time and in vivo. The recent FDA approval given to hyperpolarized [1-13C]pyrvuate for clinical studies of prostate cancer suggests that hyperpolarized 13C MR methods may be available for human studies of cardiovascular disease in the near future. While the clinical applications of cardiac hyperpolarized 13C MR remain speculative, the technique has potential to 1) advance basic knowledge, 2) improve diagnosis, and 3) optimize treatment, of ubiquitous conditions, such as myocardial ischemia and heart failure, as well as rare diseases, such as metabolic cardiomyopathies (Table 1). Future work with hyperpolarized 13C MR should focus on hardware and software development, data quantification, and development of new, highly polarized, tracers to facilitate translation of the technology into the clinic and its application in clinical studies. It is possible that, in the future, hyperpolarized methods will form an important part of routine clinical assessment in cardiology.
For their constant hard work, enthusiasm and support, we thank all the members of the Cardiac Metabolism Research Group, particularly Prof. Sir George Radda, Dr Helen Atherton, Dr Lowri Cochlin, Mr Daniel Ball and Mr Michael Dodd. We also thank Prof. Lionel Opie, Dr. Chuck Cunningham, and Dr. Albert Chen for engaging in many helpful discussions.
We would like to gratefully acknowledge the support of the British Heart Foundation (KC – Programme Grant RG/07/004/22659, SN – Programme Grant RG/10/002/28187, DT – Intermediate Research Fellowship FS/10/002/28078 & Project Grant PG/07/070/23365, the Medical Research Council (DT - New Investigator Award G0601490), the Wellcome Trust (MS was supported by a Sir Henry Wellcome Post-doctoral Fellowship), the Oxford BHF Centre for Research Excellence and the Oxford NIHR Biomedical Research Centre.
Subject codes:  CT and MRI,  Cardiovascular Imaging Agents/Techniques,  Biochemistry and Metabolism
Conflict of Interest Disclosures
KC: Grant support from GE Healthcare
DT: Grant support from GE Healthcare and Oxford Instruments Molecular Biotools
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